This invention relates to silicone rubber membranes and, more particularly, to blood-compatible gas-permeable silicone rubber membranes for use in a membrane artificial lung.
Clinical use of membrane artificial lungs for extra-corporeal blood gas exchange has been known since 1956. Largely due to the fact that the early versions of the membrane artificial lung were both cumbersome and unreliable, bubble or disc oxygenators gained a wider clinical acceptance. However, experience has indicated that membrane artificial lungs are potentially more suitable in extended applications, as for prolonged respiratory assistance, since they avoid the blood-gas interface associated with bubble or disc oxygenators and its attendant complications. For example, fat emboli, microembolic insults, and damage to blood proteins and lipoproteins appear to be the result of prolonged perfusion using blood-gas interface oxygenator systems, and up to the present time there has been no report of clinical or laboratory use of bubble or disc oxygenators beyond 24 hours with survival. With membrane artificial lungs, on the other hand, there have been reports of long-term (up to 2 weeks) animal perfusions and a number of long-term (up to 10 days) extracorporeal perfusions in man. Furthermore, evidence continues to accumulate showing the superiority of the membrane artificial lung over bubble or disc oxygenators even in short-term open heart surgery.
Despite these optimistic reports and continued improvement in membrane artificial lung design, certain problem areas still remain standing in the way of wider clinical acceptance of the membrane artificial lung. Most of these problems appear to be associated with the material constituting the membrane through which the blood gas exchange is effected. Such membrane must have the proper combination of blood compatibility, gas permeability, and tensile and tear strength, in order for the membrane artificial lung to be reliably effective. The most acceptable material thus far found for this purpose has been silicone rubber, which, when compounded with silica filler to increase tensile strength, can be cast into a thin membrane having excellent gas transfer properties, adequate strength, and at least a comparatively acceptable degree of blood compatibility. However, the blood compatibility of such membrane still leaves much to be desired, since the use of such membrane in a membrane artificial lung device necessitates maintenance anticoagulant treatment, for example, with heparin, during the perfusion in order to prevent thrombosis from occuring within the device, which treatment must be finely balanced in order to also avoid bleeding. Moreover, in addition to the clotting problem, this membrane also produces other adverse effects on the blood being perfused, such as transient leukopenia and severe granulocytopenia during initial minutes of perfusion, as well as a decrease in blood platelet count during perfusion, the mechanism and means of prevention of these effects as yet being unknown.
Although it has been speculated that the incorporation of silica filler in the above-described silicone rubber membrane may detract from the inherent blood compatibility of the silicone rubber, filler-free silicone rubber has a very low tensile strength and would not by itself have the required strength for use in the membrane artificial lung. Moreover, other fillers commonly employed with silicone rubber, such as, for example, carbon black, would not impart sufficient strength to silicone rubber membranes sufficiently thin to have the required gas permeability for use in the membrane artificial lung.
A previous attempt by the present inventor and co-workers to improve the blood compatibility of silicone rubber membranes, described in Trans. Amer. Soc. Artif. Intern. Organs, Volume 20A, Pages 269-276, 1974, involved the utilization of a double layer casting technique to make a composite membrane in which the first layer consisted of silicone rubber compounded with silica filler and the second layer, forming the blood-contacting surface, consisted of filler-free silicone rubber. When tested in a membrane artificial lung device, this composite membrane exhibited a somewhat higher degree of blood compatibility as compared with the single-layer silica-filled silicone rubber membrane, but still required maintenance anticoagulant treatment during perfusion in order to prevent thrombosis from occurring within the device, and still produced transient leokopenia and granulocytopenia during the initial minutes of perfusion. Thus, while the composite membrane required only one-half the heparin dose normally required with the single-layer silica-filled silicone rubber membrane, when no maintenance heparin treatment was administered clot formation began to occur with the composite membrane after approximately 18 hours of perfusion and progressed with time. In regard to blood platelet count, with no maintenance heparin treatment being administered during perfusion, the composite membrane produced no significant decrease at the initiation of perfusion, a decrease to 85% of baseline values after 6 hours of perfusion, and a decrease to 50% of baseline values after 18 hours of perfusion, in comparison with the single-layer silica-filled silicone rubber membrane which produced a decrease to 75% of baseline values after 5 minutes of perfusion, a decrease to 50% of baseline values after 30 minutes of perfusion, and a decrease to 25% of baseline values after 18 hours of perfusion. Thus, it can be seen that while this previously proposed composite silicone rubber membrane is somewhat more blood compatible than the single-layer silica-filled silicone rubber membrane, it still presents many of the same problems and leaves substantial room for improvement.